1. Field of the Invention
The present invention relates generally to manufacturing methods and systems (collectively referred to as xe2x80x9cprocessesxe2x80x9d) for the freeform shaping of calcium-containing powders. This invention more specifically relates to processes for shaping bone implants from various calcium phosphate powders and polymer-emulsion binders. Certain embodiments of these processes focus on the use of a Selective Laser Sintering(trademark) (xe2x80x9cSLS(trademark)xe2x80x9d) device to automatically and selectively fuse the polymer binder. In such processes, complex three-dimensional objects may be built by selectively fusing successive thin layers of the powdered material.
2. Description of the Related Art
Much attention has been given in the art to the development of materials to assist in the regeneration of bone defects and injuries. In 1926, DeJong observed the similarities between the powder X-ray diffraction pattern of the in vivo mineral and the hydroxyapatite (Ca5(OH)(PO4)3, xe2x80x9cCHAxe2x80x9d). Calcium compounds, including calcium sulfate (Nielson, 1944) , calcium hydroxide (Peltier, 1957), and tricalcium phosphate (xe2x80x9cTCPxe2x80x9d) (Albee et al., 1920), have been observed to stimulate new bone growth when implanted or injected into bone cavities (Hulbert et al., 1983). These materials also exhibit good biocompatibility and compositional similarities to human bone and tooth and can serve as resorbable or non-resorbable implants depending on their degree of microporosity.
Some TCP implants are known to be readily resorbable. For example, sintered TCP plugs with pore sizes between 100-200 microns have been implanted in rats (Bhashar et al., 1971). Very rapid bone formation was reportedly observed at three days after implantation, and highly cellular tissue, consisting of osteoblastic and fibroblastic proliferation, was found within the pores. At one week, the size of the implant was reduced, and new bone formation was extensive. After two weeks, connective tissue had infiltrated throughout the ceramic. During the next four weeks, the boney material within the ceramic continued to mature. Electron micrographs indicated that within clastlike cells, ceramic could be depicted in membrane-bound vesicles. The authors concluded that TCP implants were biodegradable, via phagocytosis, the ceramic did not elicit a marked inflammatory response, and connective tissue grew rapidly within the pores.
Similar results have also been reported by Cutright et al. (1972) who also implanted TCP in rat tibiae. In this study, the ceramic cavities were filled with osteoid and bone after 21 days and the TCP implant was no longer detectable after 48 days.
Larger implants in dogs are reported to elicit slower responses. Cameron et al. (1977) found that TCP implants in dog femurs were completely infiltrated with new bone by four weeks. However, after six weeks, the rate of new bone growth had slowed as the TCP was resorbed. Additionally, only 15% of a 2 cmxc3x972 cm iliac TCP implant in dogs was resorbed after 18 months (Ferraro et al., 1979).
Koster et al. (1976) reported the testing of the calcium phosphate formulations monocalcium phosphate, dicalcium phosphate, tricalcium phosphate, tetracalcium phosphate, and combinations consisting of 20% monocalcium phosphate and 80% of either di-, tri- or tetracalcium phosphate as implant materials in dog tibiae. These investigators tested both dense ceramics and porous ceramics with pore sizes between 800-1000 microns. They reported that tissue compatibility is dependent on the CaO/P2O5 ratio. All materials with ratios between 2/1 and 4/1 are compatible with the optimum ratio being about 3/1 for TCP. After 10 months, Koster et al. (1977) found that tetracalcium phosphate was resorbed only to a minor extent, but that TCP demonstrated lamellar bone growth throughout its pores. Both were found to be tissue compatible. The authors stated that the 3/1 material was not as strong as the 4/1 material and suggested that TCP should be used only in low stress areas while tetracalcium phosphate could be used in high stress environments.
Jarcho et al. (1976, 1977) reported the development of a process for preparing dense, polycrystalline, calcium hydroxyapatite (CHA), with the empirical formula 2(Ca5(PO4)3OH) or (3Ca3(PO4)2)Ca(OH)2. In this study, plugs were fabricated at 100% density and implanted in dogs. No evidence of tissue inflammation occurred, and in contrast to the porous TCP implants described above, little resorption or biodegradation was observed after six months.
Holmes (1979) reported that resorption did occur in porous CHA structures. These results led deGroot (1980) to suggest that all calcium phosphates are degradable (resorbable), but the rate is determined by the degree of microporosity. A dense calcium phosphate with negligible porosity would thus degrade only nominally. These results seem to be verified by Farris et al. (U.S. Pat. No. 4,673,355), who claim biocompatible materials with good properties over the range of Ca/P atomic, or molar, ratios from 0.1 to 1.34. (All patents and patent applications cited herein are incorporated by reference.) These ratios convert to CaO/P2O5 ratios between 0.2 and 2.68, lower than the 3.0 ratio suggested above. They suggest that the Ca/P or CaO/P2O5 ratio is not critical for implant applications. Ca/P ratios in the range 0.1 to 2.0 probably show satisfactory biocompatibility. Capano (1987) found that a Ca/P ratio of 0.5, which corresponds to calcium metaphosphate (xe2x80x9cCMPxe2x80x9d), has the best biocompatibility when implanted in small animals.
As the apatites are nearly identical in properties and chemical compositions to bone and tooth enamel, a considerable amount of synthetic effort has been done in this area. Patents in this area include: U.S. Pat. No. 4,046,858; U.S. Pat. No. 4 274,879; U.S. Pat. No. 4,330,514; U.S. Pat. No. 4,324,772; U.S. Pat. No. 4,048,300; U.S. Pat. No. 4,097,935; U.S. Pat. No. 4,207,306; and U.S. Pat. No. 3,379,541.
Several patents describe methods for treating apatite materials to render implantable shapes. These methods of heating and compaction under pressure in molds produce solid porous articles in various shapes. These patents include: U.S. Pat. No. 4,673,355; U.S. Pat. No. 4,308,064; U.S. Pat. No. 4,113,500; U.S. Pat. No. 4,222,128; U.S. Pat. No. 4,135,935; U.S. Pat. No. 4,149,893; and U.S. Pat. No. 3,913,229.
Several patents speak to the use of laser radiation to bond apatite materials to tooth and other surfaces, for example, U.S. Pat. No. 4,673,355 and U.S. Pat. No. 4,224,072.
Other patents describe the use of particulate or compacted apatite in conjunction with various compounds, filler, and cements, for example, U.S. Pat. No. 4,673,355; U.S. Pat. No. 4,230,455; U.S. Pat. No. 4,223,412; and U.S. Pat. No. 4,131,597.
The above discussion indicates that calcium phosphates or compounds, such as CHA that are substantially TCP (Monsanto, for example, markets CHA as TCP), are useful for a variety of bioceramic applications because they are biocompatible and can be fabricated into shapes that have a desirable combination of strength, porosity, and longevity for particular sorbable and non-sorbable needs.
Virtually any calcium and phosphate source can be used to prepare materials of interest. An important issue is the ratio of Ca to P or, as it is usually expressed, CaO to P2O5, molar ratio in the reactant mixture. For example, one can prepare monocalcium orthophosphate monohydrate from the reaction of CaO with orthophosphoric acid, H3PO4, as shown in equation 1: 
One could also react CHA with H3PO4 to achieve the same product, as shown in equation 2: 
Heating the orthophosphate hydrate can lead to a variety of known products, depending on the firing temperature used, as shown in equations 3-8: 
The xcex1-, xcex2-, and xcex4-forms of calcium metaphosphate are different crystal structures of the same chemical compound that happen to be stable at different temperatures. Tricalcium phosphates can be easily obtained from CHA by simply lowering the Ca/P ratio, as shown in equation 9: 
According to McIntosh et al. (1956), the orthophosphate hydrate can be converted to two crystalline forms by heating, as shown in equations 10-11: 
Similar reaction schemes can be written for producing di-calcium and tetra-calcium phosphates from CHA or any other calcium source by reacting with orthophosphoric acid or any other P2O5 source. The chemical and crystalline forms of the final product are simply set by the Ca/P or CaO/P2O5 molar ratio and the final temperature.
Five calcium phosphates which exhibit different x-ray diffraction patterns are known to be precipitated from aqueous solution at normal pressure (Van Wazer, 1958). These are Ca(H2PO4)2, Ca(H2PO4)2.H2O, CaHPO4, CaHPO4.2H2O, and crystalline precipitate of variable composition of hydroxyapatite with the base formula Ca5(OH)(PO4)3. Various forms of calcium phosphate compounds, Ca/P ratio range from 0.5 to 1, are prepared from the reaction of calcium hydroxyapatite with phosphoric acid.
Thermally dehydrated calcium phosphates are known to form a CaO and P2O5 binary system. For the CaO and P2O5 binary system, the chain phosphates appear between the orthophosphate (mole ratio of CaO/P2O5 of about 3) and metaphosphate (mole ratio of CaO/P2O5 of about 1) or ultraphosphate (mole ratio of CaO/P2O5 of less than 1). The metaphosphates, in particular, generally exhibit very high degrees of polymerization and good mechanical properties. In this binary system, with a mole ratio of CaO/P2O5 less than 55/45, a glass-like structure forms from the melt which has mechanical properties similar to those of natural teeth (Yoshihiro, 1975).
Many studies and methods, from powder compaction sintering to hot isostatic pressing, have been reported for the fabrication of CHA implants. However, sintered CHA materials by conventional techniques are generally as weak as sea coral even at high compacting pressure, because CHA decomposes at temperatures lower than the required temperature for sintering.
Some more recent advances are the development of hydroxyapatite (CHA) and calcium phosphate powders that can be processed to yield porous resorbable bone facsimiles (U.S. Pat. No. 4,673,355); the development of the SLS(trademark) process for directly shaping complex porous structures from thermally fusible polymer/ceramic powders without molds (U.S. Pat. No. 5,076,869); the development of low temperature infiltration and cementing techniques to prepare and replace the polymer binder with ceramic binder (U.S. Pat. No. 5,284,695); and the development of techniques for converting computed tomographic (xe2x80x9cCTxe2x80x9d) information into three-dimensional mathematical files that can automatically guide the SLS(trademark) process (Levy et al., 1992; Levy et al., 1994).
More recent work has been directed at expanding the utility of the SLS(trademark) apparatus by preparing polymer-coated ceramic powders from spray dried mixtures of water, inorganic particulate, and a custom-synthesized, emulsified, nanometer-sized, polymer binder (Barlow, 1992; Vail et al., 1992). Ceramic composites made by this approach are relatively large, 10-50 microns, agglomerates of polymer-coated inorganic particles. These agglomerate powders may spread easily into uniform layers and fuse readily in the SLS(trademark) machine to yield porous xe2x80x9cgreenxe2x80x9d parts that have relative densities near 50%, excellent connected internal porosity, and sufficient strengths to be easily handled and shipped. Interconnected pores in bioceramics are often difficult to achieve and are very important in fostering bone growth and for preparing metal matrix/ceramic parts, e.g., artificial hips.
Polymethyl methacrylate (PMMA) has also been used to form green composites with alumina and with silica/zircon (U.S. Pat. No. 5,284,695). In this process, an appropriate ceramic silicate colloid is used to infiltrate the connected pores of the polymer-bound green part, the colloid is solidified below the fusion temperature of the binder to maintain part geometry, the binder is then thermally removed and the part fired at typically 1000xc2x0 C. to form porous, all ceramic parts that are suitable for use as cores and molds for metal castings. Such parts typically have only a 1% linear shrinkage, relative to the green state. Their strengths and porosities can be adjusted by additional infiltration and firing treatments.
Lagow and co-workers have recently described the chemical synthesis of high strength CHA (U.S. Pat. No. 4,673,355) and long-chain calcium polyphosphate bioceramic powders (xe2x80x9cCPBxe2x80x9d) (Capano, 1987; Nelson et al., 1993). CPB powder is a pure calcium phosphate material with condensed phosphate chains (as shown below) with degrees of polymerization often greater than 120. 
These materials produce sintered materials that have compressive strengths greater than 200,000 psi and flexural strengths in excess of 20,000 psi. These strengths are about twice that of porcelain used to make dental crowns. Using the Lagow CHA material, Lagow and Friedman have recently completed the first successful, year duration, mandible implant in a canine. Work with CPB implants has demonstrated by electron microscopy backscattering that new bone growth occupied nearly 55% of the volume of a CPB implant in the alveolar (tooth bearing) ridge of a dog, after only four months (Nelson et al., 1993). This rate of resorption and replacement by living bone in CPB is about twice as fast as that in CHA.
The lack of suitable bone replacement is a general problem that can be potentially solved by the development of synthetic bones and bone templates that are converted to bone by the body. Bone banks currently provide gamma radiation-treated cadaver bones for various orthopedic and reconstructive purposes in a world-wide business. Appropriate geometries are not always available from these sources, and there is some concern about the transmission of HIV and other diseases. For example, in connection with spinal fusions, there is a substantial need for wedge materials that can provide support and promote the deposition of additional bone. These needs could be rapidly multiplied, provided viable materials and processes could be developed to readily provide bone materials that are shaped to the needs of each individual patient.
Facial and cranial reconstructive surgery is an area where the need for individual implant geometries is especially critical. At present, such reconstructions tend to be very difficult surgical procedures, typically involving highly skilled grafting with allogenic bone. The method and system of the present invention can be utilized to accurately construct a complete facsimile bone structure, suitable for implantation, employing geometric information that is obtained from either CT data or a Computer Aided Design (xe2x80x9cCADxe2x80x9d) software package.
The present invention addresses the foregoing and other problems experienced in the art by providing processes for automatically shaping bone implants from various calcium phosphate powders. This invention employs polymeric binder compositions particularly adapted for the formulation of free flowing calcium phosphate/binder composite powders. The powders are suitable for production of bioceramic computer-modeled geometrical implants. Green parts produced in this manner may be post processed to be substantially free of the binder. This invention also allows bioceramic parts to be produced by the low power lasers used in the laser sintering process. The processes may utilize a process called Selective Laser Sintering(trademark) in which complex three-dimensional objects can be built automatically by selectively fusing successive thin layers of powdered material.
One embodiment of the present invention provides a method for making an implant by forming a mixture of a calcium phosphate and a polymer binder, and selectively fusing the polymer binder to form an implant. These steps can be repeated to prepare a multiple-layered implant by successively forming layers of the calcium phosphate and polymer binder mixture, and selectively fusing the polymer binder in that layer and to other adjacent layers to form a plurality of connected layers.
As used herein the term xe2x80x9cimplantxe2x80x9d refers to a device that is fabricated for the purpose of embedding, or placing, within a body. The types of implants encompassed by the present invention include implants suitable for the replacement, repair, or modification of bones, teeth, and the like. However, under certain circumstances it may be conceivable that implants of the present invention may serve other useful purposes.
As used herein the term xe2x80x9cselectively fusingxe2x80x9d refers to the process of selectively coalescing, or combining, particles such that the formed structure has sufficient strength to be handled and further processed, as desired. The term xe2x80x9cselectivelyxe2x80x9d is used to denote the controlled and discriminating fashion with which the fusing process occurs. In this aspect of the present invention, xe2x80x9cfusingxe2x80x9d refers to the viscous sintering of polymer binder particles that are coating, or otherwise associating with, calcium phosphate particles. This results in a linking of the calcium phosphate particles into a part, which can be further processed by thermally decomposing and removing the polymer binder or by infiltration and subsequent thermal dehydration and the like. This fusing can generally be accomplished selectively by controlling the spatial arrangement of the interconnected particles, for example, with laser sintering processes and the like. Alternatively, the selective fusing could be accomplished using a thermal mask system or by the selective spraying of liquid binders and solvents.
The thickness of the layers formed in this method is preferably from about 3 to about 12 thousandths of an inch. In cases where CT data is employed to shape the implant, the thickness of the layers may be determined by the CT data.
The calcium phosphate is preferably prepared by reacting a mixture of hydroxyapatite and phosphoric acid, although other calcium phosphates can be used. Preferred calcium phosphates include calcium metaphosphate, calcium pyrophosphate, calcium phosphate with from about 25 to about 45 percent by weight calcium oxide, and calcium phosphate with from about 0.5 to about 2 percent by weight sodium oxide, with calcium metaphosphate being particularly preferred.
The calcium phosphate preferably has a mean particle size of from about 5 to 100 microns, with the range of from about 30 to 50 microns being most preferred. Smaller particles tend to produce weaker green parts whereas larger particles can affect the layer thickness and forming.
Certain embodiments of this invention involve coating the calcium phosphate particles with polymeric binders to provide free-flowing powders with advantageous properties for processing into shapes by sintering with a laser beam. The polymeric binder compositions may be employed to mix with, or to coat, ceramic particles to produce free-flowing powders with flow characteristics substantially independent of relative humidity.
The polymer binder may be selectively fused to replicate or form a desired geometrical shape, such as a bone or an enhancement of a bone feature. This desired geometrical shape may be obtained from CT data or CAD software data and communicated to a laser beam by a computer.
The calcium phosphate powder may be mixed with water and a polymer-emulsion binder to form a slurry. In a preferred embodiment, this slurry is rapidly dried by momentarily suspending drops of it in a stream of hot air at a temperature above the fusion temperature of the binder, such as exists in a spray drier or fluidized bed coater. The binder and powder preferably agglomerate and adhere together to form a free flowing composite powder with preferred dimensions in the range of 5-75 xcexcm.
Although it is preferable that the mixture be in the form of agglomerated polymer-coated calcium phosphate particles, mixtures of uncoated calcium phosphate powder and spray dried polymer binder can also be used. Coating the calcium phosphate with the binder is preferred as the polymer binder is used more efficiently in this embodiment, and the coating reduces segregation by density during storing or transporting.
As used herein the term xe2x80x9cagglomerated polymer-coated calcium phosphate particlesxe2x80x9d refers to an indiscriminately formed cluster of particles consisting of calcium phosphate powders that have been coated with a polymer binder. These clusters may be free-flowing substantially independent of the relative humidity.
The implant may be thermally sintered. This can effectively remove the polymer binder, and sinter, or fuse, the calcium phosphate powder.
As used in this aspect of the invention, the terms xe2x80x9csinterxe2x80x9d and xe2x80x9csinteringxe2x80x9d refer to the forming of a coherent bonded mass by heating without melting. In the case of post-processing, the calcium phosphate particles may be combined into a coherent mass with heating, whereas in sintering by laser the polymer binders may be selectively fused by the low energy of the laser beam employed.
Alternatively or in addition to thermal sintering, the implant may be infiltrated with a calcium phosphate solution or the like. This may decrease shrinkage in the implant and also modify the relative density, porosity, and other properties of the implant.
As used herein the term xe2x80x9cinfiltratingxe2x80x9d refers to a process in which a porous implant is placed in an aqueous solution of an inorganic material. This allows the solution to fill the interconnected pores of the implant, and thus upon drying deposit the inorganic material inside the implant. A further thermal sintering step may be undertaken to fuse or coalesce the calcium phosphate.
Preferred polymeric binders include those formed from 1,1-disubstituted vinyl monomers such as esters and amides of methacrylic acid and its derivatives. Examples of 1,1-disubstituted vinyl monomers include methacrylic acid, dimethylamino ethylmethacrylate and methacrylamide, methyl methacrylate and butyl methacrylate. The polymers formed from these monomers are particularly preferred because the major thermal decomposition route is depolymerization to gaseous products in both oxidizing and reducing atmospheres, largely eliminating problems with residual ash.
As used herein the phrase xe2x80x9chomopolymer, copolymer, or terpolymer of methyl methacrylatexe2x80x9d refers to polymers that are formed by polymerizing methyl methacrylate. These polymers may be formed by the homopolymerization of methyl methacrylate or by polymerizing methyl methacrylate with one or more other monomers.
Another embodiment of the present invention encompasses the implants Produced by the previously described methods. These implants preferably have a mean pore size of from about 50 to about 300 microns and a percent relative density of from about 50 to about 80%. As used herein the term xe2x80x9cpercent relative densityxe2x80x9d refers to the ratio of the implant density to the calcium phosphate density multiplied by one hundred. As defined, the percent relative density can be used to obtain the percent porosity by subtracting the percent relative density from one hundred.
Another embodiment of the present invention provides a system for making an implant, comprising a mixture of a calcium phosphate and a polymer binder, and means for selectively fusing the polymer binder to form an implant. The fusing means may comprise a laser sintering machine. The system may additionally comprise means for controlling the fusing means in order to form an implant having a desired geometrical shape. This controlling means may be a computer that obtains information about the desired geometrical shape from patient computed tomographic data or Computer Aided Design software.